Method for continuously monitoring cardiac output

ABSTRACT

A method and apparatus for continuously monitoring cardiac output based upon the phase shift between an input signal and a temperature signal indicative of the change in temperature of blood leaving the heart. In a first preferred embodiment of a cardiac output monitoring system (10), a catheter (14) is provided with an electrical resistance heater (22). An electrical current having a sinusoidal wave form with a period of from 30 to 60 seconds is applied to the heater, causing power to be dissipated into the blood within a patient&#39;s heart (12). A temperature sensor (24) disposed near a distal end of the catheter produces a signal indicative of the temperature of blood leaving the heart. The temperature signal and the signal corresponding to the electrical power dissipated in the heater (an input signal) are filtered at a frequency ω corresponding to the frequency of the applied electrical current, i.e., the frequency of the input signal. The amplitude of the input power, the amplitude of the temperature signal, and their phase difference are used in calculating cardiac output. In another embodiment, a temperature conditioned saline solution (84 ) is circulated through a catheter (14&#39;) in a closed loop, so that it flows through a heat exchanger (60) disposed within the heart. The fluid is circulated through the catheter on a periodic basis, providing a periodic input signal. The temperature signal produced by the temperature sensor on the catheter distal end and power dissipated to or absorbed from the blood by the heat exchanger comprise the two signals from which the cardiac output is determined as described above. The determination of cardiac output is also corrected for the time constant of the catheter/heater (or heat exchanger) and of the temperature sensor.

This is a divisional of the prior application Ser. No. 07/815,068 filedDec. 27, 1991, now U.S. Pat. No. 5,217,019, the benefit of the filingdate of which is hereby claimed under 35 U.S.C. § 120.

FIELD OF THE INVENTION

This invention generally pertains to apparatus and a method formonitoring the volumetric output of a heart, and more specifically, toapparatus and a method for making this determination by using aninjectateless technique that changes the temperature of blood in theheart.

BACKGROUND OF THE INVENTION

Cardiac output, the volumetric rate at which blood is pumped through theheart, is most often determined clinically by injecting a bolus ofchilled saline or glucose solution into the right auricle or rightventricle through a catheter. A thermistor disposed in the pulmonaryartery is used to determine a temperature-time washout curve as thechilled injectate/blood mixture is pumped from the heart. The area underthis curve provides an indication of cardiac output. Although thisthermo-dilution method can give an indication of cardiac output at thetime the procedure is performed, it cannot be used for continuouslymonitoring cardiac output. Moreover, the frequency with which theprocedure is performed is limited by its adverse effects on a patient,including the dilution of the patient's blood that occurs each time thechilled fluid is injected. In addition, the procedure poses an infectionhazard to medical staff from blood contact, and to the patient, fromexposure to possibly contaminated injectate fluid or syringes.

Alternatively, blood in the heart can be chilled or heated in aninjectateless method by a heat transfer process using atemperature-conditioned fluid that is pumped in a closed loop, towardthe heart through one lumen within the catheter and back through anotherlumen. The principal advantages of using such a non-injectate heattransfer process to change the temperature of blood are that the bloodis not diluted, and the temperature differential between the blood andthe heat exchanger is much less compared to the temperature differentialbetween an injectate fluid and blood in the typical thermal dilutionprocedure.

U.S. Pat. No. 4,819,655 (Webler) discloses an injectateless method andapparatus for determining cardiac output. In Webler's preferredembodiment, a saline solution is chilled by a refrigeration system orice bath and introduced into a catheter that has been inserted through apatient's cardiovascular system into the heart. The catheter extendsthrough the right auricle and right ventricle and its distal end isdisposed just outside the heart in the pulmonary artery. A pump forcesthe chilled saline solution through a closed loop fluid path defined bytwo lumens in the catheter, so that heat transfer occurs between thesolution and blood within the heart through the walls of the catheter. Athermistor disposed at the distal end of the catheter monitors thetemperature of blood leaving the heart, both before the chilled fluid iscirculated through the catheter to define a baseline temperature, andafter the temperature change in the blood due to heat transfer with thechilled saline solution has stabilized. Temperature sensors are alsoprovided to monitor both the temperature of the chilled saline solutionat or near the point where it enters the catheter (outside the patient'sbody) and the temperature of the fluid returning from the heart. Inaddition, the rate at which the chilled solution flows through thecatheter is either measured or controlled to maintain it at a constantvalue. Cardiac output (CO) is then determined from the followingequation: ##EQU1## where V_(I) equals the rate at which the chilledfluid is circulated through the catheter; ΔT_(I) equals the differencebetween the temperature of the chilled fluid input to the catheter andthe temperature of the fluid returning from the heart; ΔT_(B) equals thedifference between the temperature of the blood leaving the heart beforethe chilled fluid is circulated and the temperature of the blood leavingthe heart after the chilled fluid is circulated (after the temperaturestabilizes); and C is a constant dependent upon the blood and fluidproperties. The patent also teaches that the fluid may instead be heatedso that it transfers heat to the blood flowing through the heart ratherthan chilled to absorb heat.

U.S. Pat. No. 4,819,655 further teaches that the cardiac monitoringsystem induces temperature variations in the pulmonary artery that arerelated to the patient's respiratory cycle and are therefore periodic atthe respiratory rate. Accordingly, Webler suggests that the signalindicative of T_(B) ' (the temperature of the chilled blood exiting theheart) should be processed through a Fourier transform to yield a periodand amplitude for the respiratory cycle, the period or multiples of itthen being used as the interval over which to process the data todetermine cardiac output.

Another problem recognized by Webler is the delay between the times atwhich circulation of the chilled fluid begins and the temperature of theblood in the pulmonary artery reaches equilibrium, which is caused bythe volume of blood surrounding the catheter in the right ventricle andin other portions of the heart. The patent suggests introducing agenerally corresponding delay between the time that temperaturemeasurements are made of the blood before the chilled fluid iscirculated and after, for example, by waiting for the ΔT_(B) value toexceed a level above that induced by respiratory variations. However,for a relatively large volume heart and/or very low cardiac output, theT_(B) ' data do not reach equilibrium in any reasonable period of time.The quantity of blood flowing through the large volume heart representstoo much mixing volume to accommodate the technique taught by Webler forprocessing the data to determine cardiac output. As a result, themeasurement period for equilibrium must be excessively long to reachequilibrium, thereby introducing a potential error in the result due toeither a shift in the baseline temperature of the blood or changes inthe cardiac output. For this reason, the technique taught by Webler todetermine cardiac output using the data developed by his system is notpractical in the case of large blood volumes in the heart and/or lowcardiac outputs.

The technique disclosed by Webler also assumes that all of the energyabsorbed by a chilled fluid (or lost by a heated fluid) represents heattransferred between the fluid and the blood in the heart. Thisassumption ignores the heat transfer that occurs between the fluid andthe mass of the catheter. A somewhat smaller source of error arises dueto the energy required to change the temperature of the small thermalmass of the thermistor bead that monitors the temperature of bloodleaving the heart. For long measurement periods, these errors cangenerally be ignored. In addition, if the thermistor bead is selected tohave a very small mass and fast response time, its error contributionmay be insignificant. However, as the measurement period becomesshorter, the effect of these error sources becomes increasingly moreimportant.

Instead of cooling (or heating) the blood in the heart by heat transferwith a circulating fluid to determine cardiac output, the blood can beheated with an electrical resistance heater that is disposed on acatheter inserted into heart. The apparatus required for this type ofinjectateless cardiac output measurement is significantly less complexthan that required for circulating a fluid through the catheter. Anelectrical current is applied to the resistor through leads in thecatheter and adjusted to develop sufficient power dissipation to producea desired temperature rise signal in the blood. However, care must betaken to avoid using a high power that might damage the blood byoverheating it. An adequate signal-to-noise ratio is instead preferablyobtained by applying the electrical current to the heater at a frequencycorresponding to that of the minimum noise generated in the circulatorysystem, i.e., in the range of 0.02 through 0.15 Hz. U.S. Pat. No.4,236,527 (Newbower et al.) describes such a system, and moreimportantly, describes a technique for processing the signals developedby the system to compensate for the above-noted effect of the mixingvolume in the heart and cardiovascular system of a patient, even onewith a relatively large heart. (Also see J. H. Philip, M. C. Long, M. D.Quinn, and R. S. Newbower. "Continuous Thermal Measurement of CardiacOutput," IEEE Transactions on Biomedical Engineering, Vol. BMI 31, No.5, May 1984.)

Newbower et al. teaches modulating the thermal energy added to the bloodat two frequencies, e.g., a fundamental frequency and its harmonic, orwith a square wave signal. Preferably, the fundamental frequency equalsthat of the minimal noise in the cardiac system. The temperature of theblood exiting the heart is monitored, producing an output signal that isfiltered at the fundamental frequency to yield conventional cardiacoutput information. The other modulation frequency is similarlymonitored and filtered at the harmonic frequency, and is used todetermine a second variable affecting the transfer function between theinjection of energy into the blood and the temperature of the blood inthe pulmonary artery. The amplitude data developed from the dualfrequency measurements allows the absolute heart output to bedetermined, thereby accounting for the variability of fluid capacity ormixing volume.

Newbower et al. does not address correcting for errors due to thethermal mass of the catheter and the thermistor bead used to monitor thetemperature of blood leaving the heart. Furthermore, the techniquetaught in Newbower et al. requires matching the dual frequency data to apredefined curve using a best fit algorithm, to determine the absolutecardiac output. Accordingly, the results are not as accurate as may bedesired, particularly in the presence of noise.

It is preferably that a non-injectate method for determining cardiacoutput be based on measured output data processed using a technique thatdoes not require fitting the output data to a curve. Cardiac outputshould also be determined by a method that compensates for the mixingvolume of the heart, regardless of its relative size, and alsocompensates for the thermal mass of the catheter and the thermistor beadused to produce the output signal. The foregoing aspects and many of theattendant advantages of the present invention will become more readilyappreciated as the same becomes better understood by reference to thefollowing detailed description, when taken in conjunction with theaccompanying drawings.

SUMMARY OF THE INVENTION

In accordance with the present invention, apparatus are provided forcontinuously monitoring a cardiac output of a heart. The apparatusinclude a catheter having a plurality of lumens that extend generallybetween a proximal end and a distal end. The distal end of the catheteris insertable into the heart through a cardiovascular system. Means arealso included for supplying a periodically varying, temperaturemodifying input signal to a portion of the catheter that is spaced apartfrom its distal end. A blood temperature sensor is disposed adjacent thedistal end of the catheter and produces a blood temperature signal thatis indicative of a temperature of blood flowing from the heart. Meansare operative to determine power transferred by the temperaturemodifying input signal, producing a corresponding periodically varyingpower signal that is indicative of the power transferred. A phasecomparator determines a difference in phase between the periodicallyvarying power signal and the periodically varying temperature signal.Control means then determine the cardiac output of the heart as afunction of the power signal, the blood temperature signal, and thedifference in phase between said signals.

Preferably, the means for supplying the periodically varying temperaturemodifying signal comprise a source of an electrical current connected bya plurality of leads to a resistor disposed on a portion of the catheterthat is spaced apart from its distal end. The input signal comprises aperiodically varying electrical current that is applied to heat theresistor and any blood around the resistor. The means for determiningpower transferred comprise means for determining the power dissipated inthe resistor by the electrical current flowing through it.

Alternatively, the means for supplying the periodically varyingtemperature modifying signal comprise a pump that delivers atemperature-conditioned fluid through a closed loop fluid flow pathdefined by the lumens in the catheter. The pump cycles on and offperiodically at a predefined frequency. For this embodiment, the meansfor determining power transferred comprise a first temperature sensorthat monitors the temperature of the temperature-conditioned fluidpumped into the catheter, a second temperature sensor that monitors thetemperature of the temperature-conditioned fluid as it returns from theheart, and means for determining the rate of flow of saidtemperature-conditioned fluid. The control means determine the powertransferred as a function of the difference in temperature of thetemperature-conditioned fluid monitored by the first and the secondtemperature sensors, and the rate of flow of the temperature-conditionedfluid in the catheter.

Instead of being chilled, the temperature-conditioned fluid may beheated substantially above a normal temperature of blood entering theheart.

The cardiac output is defined by the following equation:

    CO=|P(ω)|* COS ΔΦ/(|T(ω)|*Cb)          (2)

where:

CO=the cardiac output:

P(ω)=the power transferred by the input signal, which varies at afrequency ω:

ΔΦ=the difference in phase between the power signal and the bloodtemperature signal;

T(ω)=the blood temperature indicated by the blood temperature signal,which varies at the frequency ω; and

Cb=a specific heat times density constant for the blood.

The apparatus further comprises filter means for filtering the powersignal and the blood temperature signal to remove frequencies differentfrom the frequency at which the input signal periodically varies. Inaddition, the control means compensate for an attenuation of the bloodtemperature signal by the catheter and the blood temperature sensor, indetermining the cardiac output.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a block diagram of a first embodiment of the presentinvention, illustrating the disposition of a catheter and electricalresistance heater within a human heart that is cut away to more clearlyshow the righ auricle, ventricle, and pulmonary artery;

FIG. 2 is a cut away view of a human heart, showing the disposition of acatheter through which a temperature-conditioned fluid is circulated tochange the temperature of the blood within the heart;

FIG. 3 is a block diagram of a cardiac output measurement system used inconnection with a noninjectate system that changes the temperature ofblood in the heart by heat exchange with a fluid circulated through acatheter in a closed loop; and

FIG. 4 is a flow chart showing the logical steps used in determiningcardiac output in accordance with the present invention.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT

A first embodiment of a cardiac output monitoring system in accordancewith the present invention is shown generally in FIG. 1 at referencenumeral 10. A human heart 12 is schematically illustrated in thisfigure, with a portion of the heart cut away to show the disposition ofa catheter 14 that is inserted through a patient's cardiovascular systemand into heart 12. Catheter 14 has a proximal end 16 and a distal end18. A plurality of leads 20 extend longitudinally through catheter 14(within lumens that are not separately shown) and include leads 20a and20b that carry an electrical current to an electrical resistance heater22. In the preferred form of the invention, heater 22 comprises a coilof insulated copper, stainless steel, nickel, or nichrome wireapproximately 12 cm in length that is wound around catheter 14approximately 10 to 15 cm from distal end 18. Heater 22 has a nominalresistance of from 15 to 30 ohms. Leads 20c are connected to atemperature sensor 24, which is spaced apart from distal end 18 andgenerally mounted on the external surface of the catheter so that it canreadily sense the temperature of blood flowing past the distal end asthe blood is pumped from heart 12.

As shown clearly in FIG. 1, catheter 14 extends through a right auricle26, a right ventricle 28, and into a pulmonary artery 30 of the patientwhose cardiac output is being monitored. Adjacent distal end 18 isdisposed a balloon 32, which is inflated to float distal end 18 upwardlyfrom right ventricle 28 into pulmonary artery 30. Heater 22 can bepositioned entirely within right auricle 26, or as shown, may extendfrom right auricle 26 into right ventricle 28.

A regulated current supply 34 supplies a periodic electrical currentused to generate heat in a sinusoidal wave form at heater 22, at avoltage ranging from 10 to 25 volts peak amplitude. Alternatively, asquare wave current supply can be used. As the current flows through thewire coil comprising heater 22, it produces heat in proportion to the I²R losses in the heater (where I is the current and R is the resistanceof the heater). The heat produced is transferred to the blood withinright auricle 26 and right ventricle 28. A current sensor 36 produces asignal indicative of the magnitude of the electrical current flowingthrough lead 20a to heater 22, and this signal is input through leads 38to analog-to-digital (A-D) converters 40. A second input to A-Dconverters 40 is a voltage signal that indicates the voltage developedacross heater 22; this voltage signa is conveyed by a lead 42. The thirdinput to the A-D converters comprises the signal indicative of thetemperature of the blood leaving heart 12, produced by temperaturesensor 24, connected to leads 25, which comprise the distal end of leads20c. Digitized signals from A-D converters 40 are conveyed through leads44 to input ports (not separately shown) on a portable computer 46.

Associated with portable computer 46 is a video display 48 on which datadefining the cardiac output of heart 12 are displayed, along with otherdata and information. A keyboard 50 is connected to portable computer 46to provide for input and user control of the cardiac output measurement.In addition, portable computer 46 includes a hard drive or floppy drive52 that is used for magnetic storage of data, test results, and programssuch as the software controlling the measurement of cardiac output.Portable computer 46 controls regulated current supply 34 by supplyingcontrol signals transmitted through leads 54 that extend between theregulated current supply and the portable computer.

The electrical current that energizes heater 22 to heat the bloodflowing through heart 12 is supplied either in the form of a sine wavehaving a 30- to 60-second period, or as a square wave with an energizedperiod ranging between 15 and 30 seconds (followed by a like durationduring which no current is supplied). The power developed by heater 22thus represents a periodic input signal, whereas the signal developed bytemperature sensor 24 comprises an output signal indicative of thetemperature of the blood leaving the heart. To determine powerdissipated within heater 22, the digitized signals indicative of thecurrent flowing through the heater and voltage drop across it aremultiplied together by portable computer 46. The power dissipated withinheater 22 to heat the blood flowing through heart 12, i.e., the peak topeak amplitude, is therefore easily determined and is defined as the"input signal" for purposes of the following discussion. Accordingly,the power applied, which represents the input signal, and thetemperature of the blood exiting the heart through the pulmonary artery,which represents the output signal, are used in the first preferredembodiment to determine the cardiac output of heart 12, as explainedbelow.

An alternative embodiment for developing an input signal and an outputsignal that can be used to determine the cardiac output of heart 12 isshown in FIG. 2. In this embodiment, a catheter 14' is used to convey acooling or heating fluid to a heat exchanger 60 formed on the catheter,set back from its distal end so that the heat exchanger is within rightauricle 26. Two lumens (not separately shown) within catheter 14' definea supply fluid path 62 through which a liquid cooled to a temperaturewell below that of the body temperature of the patient is conveyed toheat exchanger 60, and a return fluid path 64 through which the fluid isthen returned back to a source of the fluid, external to the patient'sbody. In most other aspects of its configuration and use, catheter 14'is similar to catheter 14, shown in FIG. 1. Like catheter 14, catheter14' includes temperature sensor 24 disposed adjacent its distal end 18so that it is positioned within pulmonary artery 30.

Instead of cooling a fluid to a temperature lower than the temperatureof blood entering heart 12 through catheter 14', the fluid may be heatedabove the temperature of the blood so that it transfers heat to theblood, just as heater 22 does. In either case, whether the input signalcools the blood or heats it, the cardiac output measurement systemchanges the temperature of blood in the heart on a periodic basis sothat the output signal produced by temperature sensor 24 changesperiodically in response thereto. Furthermore, the change in thetemperature of blood flowing from the heart, i.e., the output signal, isphase shifted relative to the input signal due to the time required tochange the temperature of the mass of blood within the right auricle andright ventricle.

In FIG. 3, the remainder of a cardiac output measurement system 80,which is used for circulation of a temperature conditioned fluid (withrespect to the temperature of blood entering heart 12) through catheter14' is illustrated schematically. Cardiac output measurement system 80includes a reservoir 82 (hanger bag) of a saline solution 84. Salinesolution 84 flows under the influence of gravity through a line 86 to apump 88. When energized for periods of 15 to 30 seconds at a time, pump88 forces saline solution 84 through a supply line 90, which isconnected to supply fluid path 62 within catheter 14'. After the liquidflows through catheter 14' and exchanges heat with blood within heart 12at heat exchanger 60, it flows back along return fluid path 64 into areturn line 92. Return line 92 passes through an external heat exchanger96, which reduces the temperature of the returning saline solution toambient temperature, e.g., 24° C. Thereafter, the returning salinesolution flows back into reservoir 82 for recirculation by pump 88.

The operation of pump 88 is controlled by a pump control 98, which isconnected to the pump by leads 100 that convey signals determining therate at which pump 88 operates. In addition, leads 100 carry an ENABLEsignal that energizes pump 88 and signals indicative of any alarmcondition, e.g., air in the line or restriction of lines 86 or 90. Pumpcontrol 98 also receives a signal from pump 88 indicating that the pumpis running, to confirm that fluid is being delivered to the catheter asexpected.

It will be appreciated that instead of using liquid at ambienttemperature to cool the blood flowing through the heart, saline solution84 can be chilled to a much cooler temperature (using a chiller coildisposed downstream of pump 88, in heat transfer relationship withsupply line 90). For example, saline solution 84 can be chilled to alower than ambient temperature by heat transfer with ice water at 0° C.;or, a more elaborate evaporative refrigerant chiller coil can beemployed that uses a refrigerant fluid to cool saline solution 84 as therefrigerant fluid undergoes expansion. Similarly, it is also possible toprovide heat transfer between saline solution 84 that is circulatedthrough catheter 14' and a heated liquid or to provide heat from someother source so that the saline solution entering catheter 14' iselevated in temperature above the temperature of blood entering heart12.

Pump control 98 is controlled by portable computer 46 so that pump 88 isenabled on a periodic basis to circulate temperature conditioned salinesolution 84 through catheter 14'. In this embodiment, the input signalto the blood within the heart is represented by the flow of temperatureconditioning liquid through catheter 14'. A signal applied to pumpcontrol 98 over lines 103, which connect the pump control to theportable computer, is used to enable the operation of pump 88. The flowof temperature-conditioned saline solution 84 through catheter 14' isenabled for 15 to 30 seconds, then turned off for an equivalentinterval, and this cyclic operation is continued during the measurementof cardiac output.

A plurality of lines 102 carry signals indicative of varioustemperatures to A-D converters 40, which supplies the correspondingdigitized signals to portable computer 46. Specifically, a line 102a isconnected to lead 20c, and thus conveys the signal indicative of thetemperature of blood leaving heart 12 to A-D converters 40. A lead 102bis connected to a temperature sensor 104 that produces a signalindicative of the temperature of saline solution 84 flowing into supplyfluid path 62 within catheter 14'. Similarly, a temperature sensor 106is connected to a lead 102c, which conveys a signal indicative of thetemperature of saline solution 84 returning from catheter 14' intoreturn line 92. A plurality of fluid lines 94 are connected to otherlumens within catheter 14' and can be used to inject medication into theheart and inflate balloon 32 during the insertion of catheter 14' intoheart 12.

As noted in the Background of the Invention, the present inventionenables cardiac output to be determined continuously rather thanintermittently (an unfortunate limitation of the conventional injectatethermal dilution technique) and is much less prone to noise thanprevious continuous cardiac output monitoring methods. In the presentinvention, cardiac output is determined by portable computer 46following the logical steps shown in a flow chart 120, in FIG. 4.Starting at a block 122, the temperature of blood flowing through heart12 is modified by applying the input signal, e.g., by electrical currentto heater 22, or by initiating the flow of a temperature-conditionedfluid through catheter 14' so that heat is transferred at heat exchanger60--in either case, thereby modifying the temperature of blood withinthe heart. The transfer of heat to or from blood within heart 12 occursat a frequency ω, as shown in block 122. This frequency is selected tominimize the noise caused by patient respiration.

A dashed-line block 124 indicates that the blood heated or cooled by theinput signal mixes with the other blood in right ventricle 28 and enterspulmonary artery 30. A block 126 refers to temperature sensor 24, whichproduces the signal that is indicative of the temperature of bloodexiting heart 12. With reference to a block 128, the blood temperature Twithin pulmonary artery 30 comprises the output signal that is digitizedby A-D converter 40. The digitized signal indicative of the temperatureof blood within the pulmonary artery is filtered at the input frequencyω, as indicated in a block 130 in FIG. 4.

In the preferred embodiment, the output signal is filtered by portablecomputer 46. Specifically, a discrete Fourier transform is performed onthe digitized output signal to transform the signal from the time domaininto the frequency domain. The portion of the transformed signal at theinput frequency ω is thus determined and comprises a filtered outputsignal. By filtering the output signal (and the input signal, asdescribed below), noise at other frequencies is substantiallyeliminated. Alternatively, an analog bandpass filter circuit could beused to process the input signal before it is digitized, in lieu of thediscrete Fourier transform. Other types of digital filtering could alsobe used to eliminate noise components at other frequencies.

After the output signal is filtered, the amplitude of the filteredoutput signal is determined, as noted in a block 132. Portable computer46 uses the peak to peak value of the filtered output signal for thisamplitude, represented by |T(ω)|. The value |T(ω)| is then used in ablock 134 for calculating cardiac output. Since the filtered outputsignal is a periodically varying signal, it has a phase relationshipthat is represented by the value Φ_(out) (used as described below).

The left side of flow chart 120 is directed to the steps used inprocessing the input signal. As shown in a block 138, the heating orcooling power P, which represents the heat transferred to or absorbedfrom the blood in the heat, is determined. As described above, theheating power of heater 22 is determined from the product of theelectrical current flowing through it and the voltage drop across theheater, as well known to those or ordinary skill in the art.

If catheter 14' is used and heat is transferred between the circulatedsaline solution and blood flowing through heart 12, the input signal isdetermined as a function of: (a) the temperature differential betweenthe saline solution supplied to catheter 14' and that returning from thecatheter, measured at temperature sensors 104 and 106, and (b) thesaline solution flow rate provided by pump 88. In the preferred form ofthe invention shown in FIG. 3, pump 88 is set to provide a flow rate ofapproximately 1.5 liters per hour when energized. The input signal(representing input power P) is determined by portable computer 46 fromthe digitized signals indicative of the saline solution temperatures attemperature sensors 104 and 106, the flow rate of the saline solutionthrough the catheter (predefined or measured), and the specific heat ofthe saline solution, as will be apparent to those of ordinary skill inthe art.

Portable computer 46 then filters the input signal at the inputfrequency ω, as indicated in a block 140. To filter the input signal,the portable computer processes it with a discrete Fourier transform,converting it from the time domain to the frequency domain. The portionof the transformed signal at the frequency ω comprises the filteredinput signal. The filtered input signal has both a phase and amplitude.In a block 142, the amplitude of the input signal is determined and isinput to block 134 as |P(ω)|. The phase of this filtered input signal.Φ_(in), is compared to the phase of the output signal in a block 136,producing a differential phase ΔΦ, which is equal to the differencebetween Φ_(in) and Φ_(out). Portable computer 46 determines thedifferential phase and as shown in block 134, calculates cardiac output"CO" as follows:

    CO=|P(ω)|*COS ΔΦ/(|T(ω)|*Cb)          (3)

In the above equation, the value Cb is the product of specific heat anddensity of blood.

The volume of blood within right ventricle of heart 12, i.e., the mixingvolume, is estimated from the following expression: ##EQU2## where τ isthe period of the input signal. To reduce the effects of phase noise onthe determination of cardiac output, an estimation of mixing volume canbe made from Equation 4 and used in the following relationship: ##EQU3##

The estimate of mixing volume is preferably averaged over a long term(assuming that volume is relatively constant over the time during whichcardiac output is determined), yielding an average mixing volume, V,which is used in Equation 5 to determine cardiac output. The resultingdetermination of cardiac output from Equation 5 is therefore lesssensitive to phase noise, including heart rate variations.

When a heat signal is injected into the blood within heart 12, either bycooling the blood or by applying heat to it, a transport delay time isincurred before the input heat signal reaches temperature sensor 24 inthe pulmonary artery. The transport delay time adds a phase shift thatis flow rate and vessel size dependent. The phase error due to transportdelay time is defined as: ##EQU4## where L is equal to the length of thepath from the point of which the heat signal is injected into the bloodwithin the heart to the point at which the temperature sensor isdisposed (in cm), R is the vessel radius (in cm), and CO is the cardiacoutput in liters/second. For example, a typical phase shift would beapproximately 28.8° for a path 10 cm in length, with a rate of flow ofone liter per minute, a radius of 1.6 cm and a period for the injectionof the heat signal equal to 60 seconds.

The phase shift introduced by transport delay becomes significant atrelatively low flow rates, making accurate correction for the mixingvolume difficult. One way to address this problem is to apply the inputsignal at two (or more) different frequencies, enabling a separateestimate of transport delay phase shift and mixing volume phase shift tobe determined from the difference in phase shift at the differentfrequencies.

There are two additional sources of error for which corrections can beapplied in determining cardiac output. The sources of error relate tothe time constant for the catheter and thermistor caused by theirrespective thermal masses. The thermal mass of the catheter attenuatesand phase shifts the input signal, whereas the thermal mass oftemperature sensor 24 attenuates and phase shifts the receivedtemperature signal corresponding to the change in temperature in theblood flowing past temperature sensor 24. The correction used in thepreferred embodiment assumes a simple first-order system. For example,heater 22 is assumed to have a time constant T_(htr) (actually the timeconstant is for the catheter and heater), and temperature sensor 24 tohave a time constant T_(sens), both of which are empirically determined.Cardiac output is then determined from: ##EQU5## where: Φ_(htr)=-ARCTAN(ω*T_(htr));

Φ_(sens) =-ARCTAN(ω*T_(sens));

HTR_(atten) =COS (Φ_(htr)); and

SENSOR_(atten) = COS (Φ_(sens)).

Equation 7 recognizes that a time delay occurs between the arrival attemperature sensor 24 of blood having a different temperature due to theinput of a heat signal and the change in the output signal of thetemperature sensor. Similarly, the thermal mass of the catheter/heaterintroduces a time delay between the application of the input signal andthe transfer of energy into the blood around heater 22 (or heatexchanger 60). Typical time constants for both heater 22 and temperaturesensor 24 are approximately two seconds each. Based on the assumptionthat the time constants for these two elements do not vary with flowrate, amplitude errors and thus cardiac output errors introduced fromthis source of error, should be constant, dependent only on thefrequency of the input signal. Accordingly, the phase shift introducedby these time constants should also be constant. Since the sensitivityto phase errors increases at low flow rates and large mixing volumes, itis important to correct for the phase shift due to the time constants ofthe catheter/heater (or heat exchanger) and temperature sensor, at largeoverall phase angles.

While the preferred embodiment of the invention has been illustrated anddescribed, it will be appreciated that various changes can be madetherein without departing from the spirit and scope of the invention.Accordingly, it is not intended that the scope of the present inventionbe in any way limited by the disclosure of the preferred embodiment, butinstead that it be determined entirely by reference to the claims thatfollow.

The embodiments of the invention in which an exclusive property or privilege is claimed are defined as follows:
 1. A method for determining a cardiac output of a heart, comprising the steps of:(a) changing a temperature of blood within the heart so that it varies periodically; (b) sensing a temperature of blood leaving the heart, producing a blood temperature signal that varies periodically; (c) producing a power signal indicative of power used to change the temperature of the blood within the heart, said power signal varying periodically; (d) determining a difference in phase between the power signal and the blood temperature signal; and (e) determining the cardiac output of the heart as a function of the power signal, the blood temperature signal, and the difference in phase between said signals.
 2. The method of claim 1, wherein the step of changing the temperature of the blood within the heart comprises the step of heating the blood with a resistance heater periodically energized with an electrical current.
 3. The method of claim 1, wherein the step of changing the temperature of the blood within the heart comprises the step of periodically circulating a temperature-conditioned fluid in heat transfer relationship with the blood in the heart.
 4. The method of claim 3, wherein the fluid is chilled to a temperature substantially below a temperature of blood entering the heart.
 5. The method of claim 3, wherein the step of producing the power signal comprises the steps of:(a) monitoring a temperature of the temperature-conditioned fluid before heat is transferred between it and the blood within the heart, to obtain a first temperature; (b) monitoring a temperature of the temperature-conditioned fluid after heat is transferred between it and the blood within the heart, to determine a second temperature; (c) determining a flow rate for the temperature-conditioned fluid; and (d) producing the power signal as a function of a difference between the first and second temperatures, and of the flow rate of the temperature-conditioned fluid.
 6. The method of claim 1, wherein the step of producing the power signal comprises the step of producing a signal that is indicative of an electrical current used in changing the temperature of the blood.
 7. The method of claim 2, further comprising the step of varying the electrical current, the electrical current being applied to sinusoidally heat a resistor that is disposed on a portion of a catheter positionable within the heart, heat dissipated by the resistor elevating the temperature of the blood in the heart so that its temperature varies sinusoidally.
 8. The method of claim 1, wherein the cardiac output is defined by an equation as follows:

    CO=|P(ω)|* COS ΔΦ/(|T(ω)|*Cb)

where: CO=the cardiac output; P(ω)=the power used to change the temperature of blood in the heart, which varies at an angular frequency ω; ΔΦ=the difference in phase between the power signal and the blood temperature signal; T(ω)=the blood temperature indicated by the blood temperature signal, which varies at the angular frequency ω; and Cb=a specific heat times density constant for the blood.
 9. The method of claim 1, further comprising the step of filtering the power signal and the blood temperature signal to filter out frequencies different from that of the power signal.
 10. The method of claim 1, further comprising the step of compensating for a temperature attenuation and a phase shift error in the blood temperature signal due at least in part to a thermal mass of a catheter inserted into the heart.
 11. The method of claim 1, further comprising the step of compensating for a temperature attenuation and a phase shift error in the blood temperature signal due at least in part to a thermal mass of a temperature sensor used to produce the blood temperature signal.
 12. The method of claim 2, wherein the step of heating the blood within the heart comprises the step of supplying the resistance heater with an electrical current that varies periodically at a plurality of frequencies, and wherein the step of determining the difference in phase between the power signal and the blood temperature signal comprises the step of determining the difference in phase of said signals at each of the plurality of frequencies.
 13. The method of claim 12, further comprising the step of separately estimating a transport delay phase shift and a mixing volume phase shift from the difference in phase of said signals at each of the plurality of frequencies.
 14. The method of claim 1, further comprising the step of estimating a mixing volume of blood within the heart, V, based on a relationship: ##EQU6## where: τ is a period of the power signal;P(ω)=the power used to change the temperature of blood in the heart, which varies at an angular frequency ω; ΔΦ=the difference in phase between the power signal and the blood temperature signal; T(ω)=the blood temperature indicated by the blood temperature signal, which varies at the angular frequency ω; and Cb=a specific heat times density constant for the blood.
 15. The method of claim 14, further comprising the step of averaging a plurality of estimates of the mixing volume of blood within the heart and determining cardiac output from an average estimate of mixing volume based on a relationship: ##EQU7## where V is the average estimate of mixing volume, use of the average estimate of mixing volume thereby reducing a sensitivity of the determination of cardiac output to phase noise, including heart rate variations. 